Ultrasound medical systems and methods include ultrasound imaging of anatomical tissue to identify tissue for medical treatment. Ultrasound may also be used to medically treat and destroy unwanted tissue by heating the tissue. Imaging is done using low-intensity ultrasound waves, while medical treatment is performed with high-intensity ultrasound waves. High-intensity ultrasound waves, when focused at a focal zone a distance away from the ultrasound source, will substantially medically affect tissue in the focal zone. However, the high-intensity ultrasound will not substantially affect patient tissue outside the focal zone, such as tissue located between the ultrasound source and the focal zone. Other treatment regimes of interest include unfocused high-intensity ultrasound wherein the ultrasound energy is distributed over a relatively broad region of tissue rather than being generally concentrated within a focal zone.
Ultrasound waves may be emitted and received by a transducer assembly. The transducer assembly may include a single transducer element, or an array of elements acting together, to image the anatomical tissue and to ultrasonically ablate identified tissue. Transducer elements may employ a concave shape or an acoustic lens to focus or otherwise direct ultrasound energy. Transducer such arrays may include planar, concave or convex elements to focus ultrasound energy. Further, array elements may be electronically or mechanically controlled to steer and focus the ultrasound waves emitted by the array to a focal zone to provide three-dimensional medical ultrasound treatment of anatomical tissue. In some treatments the transducer is placed on the surface of the tissue for imaging and/or treatment of areas within the tissue. In other treatments the transducer is surrounded with a balloon which is expanded to contact the surface of the tissue by filling the balloon with a fluid such as a saline solution to provide acoustic coupling between the transducer and the tissue.
Examples of ultrasound medical systems and methods include: deploying an end effector having an ultrasound transducer outside the body to break up kidney stones inside the body; endoscopically inserting an end effector having an ultrasound transducer into the rectum to medically destroy prostate cancer; laparoscopically inserting an end effector having an ultrasound transducer into the abdominal cavity to destroy a cancerous liver tumor; intravenously inserting a catheter end effector having an ultrasound transducer into a vein in the arm and moving the catheter to the heart to medically destroy diseased heart tissue; and interstitially inserting a needle end effector having an ultrasound transducer into the tongue to medically destroy tissue to reduce tongue volume as a treatment for snoring. Methods for guiding an end effector to the target tissue include x-rays, Magnetic Resonance Images (“MRI”) and images produced using the ultrasound transducer itself.
Low-intensity ultrasound energy may be applied to unexposed subdermal anatomical tissue for the purpose of examining the tissue. Ultrasound pulses are emitted, and returning echos are measured to determine the characteristics of the unexposed subdermal tissue. Variations in tissue structure and tissue boundaries have varying acoustic impedances, resulting in variations in the strength of ultrasound echos. A common ultrasound imaging technique is known in the art as “B-Mode” wherein either a single ultrasound transducer is articulated or an array of ultrasound transducers is moved or electronically scanned to generate a two-dimensional image of an area of tissue. The generated image is comprised of a plurality of pixels, each pixel corresponding to a portion of the tissue area being examined. The varying strength of the echos is preferably translated to a proportional pixel brightness. A cathode ray tube, computer monitor or liquid crystal display can be used to display a two-dimensional pixellated image of the tissue area being examined.
When high-intensity ultrasound energy is applied to anatomical tissue, significant beneficial physiological effects may result. For example, undesired anatomical tissue may be ablated by heating the tissue with high-intensity ultrasound energy. By focusing the ultrasound energy at one or more specific focusing zones within the tissue, thermal effects can be confined to a defined region that may be remote from the ultrasound transducer. The use of high-intensity focused ultrasound to ablate tissue presents many advantages, including: reduced patient trauma and pain; elimination of the need for some surgical incisions and stitches; reduced or obviated need for general anesthesia; reduced exposure of internal tissue; reduced risk of infection and other complications; avoidance of damage to non-targeted tissue; lack of harmful cumulative effects from the ultrasound energy on the surrounding non-target tissue; reduced treatment costs; minimal blood loss; and the ability for ultrasound treatments to be performed at non-hospital sites and/or on an out-patient basis.
Ultrasound treatment of anatomical tissue may involve the alternating use of both low-intensity imaging ultrasound and high-intensity treatment ultrasound. During such treatment, imaging is first performed to identify and locate the tissue to be treated. The identified tissue is then medically treated with high-intensity ultrasound energy, such as for the purpose of destroying the tissue. After a period of exposure to high-intensity ultrasound, a subsequent image of the tissue is generated using low-intensity ultrasound energy to determine the results of the ultrasound treatment and provide visual guidance to the user to aid in subsequent treatments. This process of applying low-energy ultrasound to assist in guiding the position and focal point of the transducer, followed by high-energy ultrasound to ablate the undesired anatomical tissue, may continue until the undesired tissue has been completely ablated.
In addition to imaging, monitoring of the temperature of the tissue being treated is desirable so that the tissue being treated, as well as the delivered treatment, can be readily visualized and controlled. For example, temperature monitoring is essential for hyperthermia treatments wherein tissue is exposed to ultrasound for the purpose of raising the temperature of the tissue. Hyperthermia treatments have been shown to be effective in the treatment of cancerous tumors, which are known to be more sensitive to heat than healthy tissue. Temperature monitoring is also useful for high-intensity ultrasound treatment of tissue, such as ablation, since the temperature of the tissue provides an indication of the extent and spatial distribution of the treatment.
The temperature of the tissue may be measured directly, such as with a thermocouple array. However, this arrangement is not desirable for several reasons. First, the thermocouple array must be placed into the region being treatment. This may require invasive placement of the array, thereby eliminating the non-invasive treatment advantages offered by ultrasound. In addition, the thermocouple array may have a relatively slow response time, providing the operator with delayed feedback regarding the status of the treatment.
Various non-invasive temperature-monitoring methods have been developed to overcome the drawbacks of thermocouple arrays. For example, magnetic resonance imaging (“MRI”) is used in the art to noninvasively measure temperature rises in vivo. However, MRI feedback requires an apparatus separate from the ultrasound treatment apparatus to provide the temperature measurement. As a result, MRI-assisted ultrasound methods are expensive and cumbersome, making such treatment impractical for routine use in surgical treatments.
The prior art has suggested using ultrasound waves to measure temperature. Previous studies have shown that temperature rises in tissue cause local thermal expansion and sound speed changes, which can be estimated from perturbations of ultrasonic echo signals. Since the travel time of an ultrasonic echo depends on both the path length (which is changed by any thermal expansion) and sound speed (which depends on temperature), temperature changes in tissue being treated cause measurable alterations to echo travel times. By estimation and further processing of these travel times, temperature rises can be approximately mapped. This arrangement is highly desirable, since heating effects of ultrasound treatment may be measured using inexpensive, portable, and unobtrusive ultrasound apparatus. For configurations in which therapy and monitoring are performed using the same ultrasound probe, this approach is particularly desirable, since a temperature map may be automatically co-registered with the therapeutic ultrasound beams.
Available approaches to pulse-echo temperature mapping have employed cross-correlation of ultrasound echo signals to estimate changes in travel time (i.e., delay) for each pixel of an image, before and after heating. These travel-time maps are then smoothed and differentiated by signal processing to obtain maps of the echo strain, which is assumed to be proportional to the temperature rise of the tissue being treated. Repetition of this process can follow incremental temperature rises associated with prolonged heating. These approaches suffer from several drawbacks, including low spatial resolution and high artifactual content. Since many time-delay estimation steps are needed to build up a temperature image, these approaches are also computationally inefficient.
There is a need for a method of measuring the temperature rise of tissue being treated by ultrasound that is noninvasive. There is a further need for method of monitoring temperature rise that is not expensive and is not cumbersome. There is a still further need for a method of monitoring temperature measurement that has relatively high spatial resolution and relatively low artifactual content, and is computationally efficient.